1. Field of the Invention
This invention relates to the field of photodetector systems, and more particularly, to the field of light intensity measuring photodetector systems.
2. Background Information
Some opto-electronic systems such as camera focus determination systems, rely on conversion of an incident optical image to corresponding electrical signals for processing, but do not require quantitative measurement of incident light intensity in order to fulfill their objective. However, such systems may require adjustment or compensation for incident light intensity to prevent saturation of system components or other effects which would prevent proper operation of the system. U.S. Pat. No. 4,437,743 to Sakai et al. discloses such a camera focus adjustment system. That system employs a CCD imager to convert optical focus information to electronic signals for processing to determine the degree of focus of the scene being viewed by the camera. The integration period for the CCD is adjusted in accordance with the incident light intensity. The system compensates for the sensor dark current variations with the length of the integration period by masking a portion of the CCD to obtain a dark current signal which is the first bit read out of the CCD. This value is sampled and held and applied to one input of a differential amplifier while the image signals are applied to the other input of the differential amplifier to provide a so-called dark current compensated signal whose amplitude is then used to adjust the CCD integration time.
Many other electronic systems require quantitative measurement of incident light intensity for their proper operation. Such systems include focal plane imaging systems, exposure setting systems for cameras, medical diagnostic systems such a photoelectric LED blood oxygen monitoring systems and computed tomography (CT) systems employing solid scintillator x-ray detection systems. For example, U.S. Pat. No. 4,781,195 to Martin discloses a blood oxygen monitoring system which measures the ambient light effect on its sensor diodes, fixes a compensation for that ambient light effect and then takes a measurement using that compensation to correct for the ambient light. The ambient light effect is then measured again and a new compensation is established for the next measurement and so forth. Thus, this system alternates measurement and compensation determination cycles. U.S. Pat. No. 3,877,039 to Ichinohe et al. discloses a camera exposure setting system which first measures an ambient light intensity by charging a capacitor and then subtracts a dark current signal by applying a current to discharge the capacitor as a means of compensating for the non-linearity introduced by its photosensitive diodes. U.S. Pat. No. 4,857,725 to Goodnough et al. discloses a diode-based focal plane imaging system which determines a compensation for the difference in the photoresponses of each associated pair of diodes and maintains that compensation over a measurement interval which is preferably an entire mission.
In a computed tomography (CT) x-ray scanning system which employs a solid, luminescent scintillator to convert incident x-rays to luminescent light a multicellular x-ray scintillation detector system is used to convert the incident x-ray intensity to electrical signals whose amplitudes are measures of the x-ray intensity. Such CT scanning systems include a gantry on which an x-ray source is mounted on one side of a measurement zone and an x-ray scintillation detector system is mounted on the opposite side of the measurement zone in alignment with the x-ray beam. Such systems preferably employ a fan-shaped x-ray beam and a scintillation x-ray detector which may comprise 1,000 or more separate detection cells disposed in a linear array.
For maximum data collection accuracy, the individual cells of the scintillator detection system are disposed immediately adjacent to each other, thereby maximizing the collection by the scintillation detector system of the x-rays emerging from the patient or other object being examined. The individual blocks of scintillator material may typically be about 1 mm wide by about 30 mm high by about 3 mm deep. The width of the scintillator block being determined by the desired data resolution in the direction of the length of the scintillation detector array, the height of the block being determined by the thickness of the x-ray fan beam in combination with the desired vertical thickness of the measurement zone and the depth of the scintillator block being determined by the x-ray stopping power of that block.
For ease of manufacture and assembly, it is preferred to have the photoresponsive diodes associated with a number of adjacent scintillator blocks integrated in a single chip or wafer of semiconductor material. The patterned side of the semiconductor structure and the electrodes thereon are disposed toward the blocks of scintillator material for maximum conversion of luminescent light to electrical signals. The immediate side-by-side placement of adjacent scintillation detector cells requires that the output connections to the individual diodes be located adjacent a narrow end of the diode in order that wires, conductors and other optical obstructions may be excluded from the active collection area of the diode to maximize light collection, system efficiency and cell-to-cell uniformity within the detector system.
To generate a computed tomography image, the gantry is rotated about the measurement zone while the x-ray beam is on and the output from the scintillation detector system is recorded or stored for concurrent or subsequent processing. Data is taken continuously during each revolution of the gantry. At each position where data is taken, the output of each of the detector cells is determined and stored.
In such a system, the vast quantity of data which has been recorded (.about.4.times.10.sup.5 data points) is subsequently processed to generate an image of the object or patient disposed in the measurement zone during the measurement process.
Each data point comprises the position of the detection cell at the time the data was taken and the amplitude of the output from the photodiode associated with that cell at that time. All of these output amplitudes are processed using computerized tomography image reconstruction techniques, which are known in the art, to generate an image of the object in the measurement zone. The accuracy of the generated image depends on the accuracy with which incident x-ray intensity on the scintillator cell is converted to an electronic signal having an amplitude which is a measure of that x-ray intensity.
Thus, image accuracy is limited by the ability of the scintillation x-ray detection system to accurately convert x-ray intensity to electronic signal amplitude. In order to maximize throughput, minimize patient x-ray exposure and inconvenience, it is considered desirable to perform a data collection scan as rapidly as possible. In order to provide accurate image generation with low intensity x-rays and high scan rates, it is necessary that the photodetection system accurately transduce incident luminescent light intensity to a corresponding electrical signal amplitude in a linear manner across a wide dynamic range including both very low and very high light intensities, a dynamic range of 10.sup.4 to 10.sup.5 or more being considered desirable.
Because of this wide desired dynamic range and the desire to be able to operate at very small rates of photo generated charge carriers (in order to minimize patient exposure to x-rays), it is necessary to minimize noise contributions to signals throughout the overall data acquisition system.
Present General Electric Company CT scanners of the Zeus type employ a solid ceramic scintillator to convert incident x-ray intensity to luminescent light. A large area PIN photosensitive diode on the order of 3/4 mm.times.30 mm in area is optically coupled to the scintillator block and detects the luminescent light produced thereby and converts that light to an output current whose amplitude is linearly related to the intensity of the luminescent light and thus to the incident x-ray intensity, provided the scintillator provides a linear x-ray-to-luminescent-light conversion. This output current is in turn converted to an output voltage by an operational amplifier which is dedicated to that photosensitive diode.
Many design considerations go into the design of the photodetector diode and its operational amplifier output system. For accurate image generation, each of the scintillator detector cells must provide substantially identical conversion of x-ray intensity to electronic signal amplitude. Consequently, it is necessary to minimize offset voltages among the operational amplifiers connected to the various photodetector diodes. This is facilitated by operating the diode in a zero bias (photoconductive) mode to minimize offset voltages and thermal drift.
Data is taken substantially continuously during a scan. That is, the output from the photosensitive diodes is continuous, as is the output from the operational amplifiers to which the diodes are connected. The output from the operational amplifiers is sampled by analog-to-digital conversion and stored at the desired data acquisition rate to provide data at each of the desired positions of the scanning gantry.
In order to produce a clear, high quality image of the object or patient being scanned, such a CT scanning system requires that the electrical output signal be directly proportional to the luminescent light intensity. Any deviation from such direct proportionality constitutes system noise which degrades the obtainable image quality. As is well known in the semiconductor art, the response characteristics of a photosensitive diode depend on the diode's temperature because of an increase in the inherent internal generation of hole-electron pairs within the semiconductor material with increasing temperature. Such internally generated thermal hole-electron pairs are indistinguishable to the overall system from the hole-electron pairs which are generated by the incident luminescent light which it is desired to measure. Thus, it is important to stabilize the diode temperature and thermal generation rate. Consequently, the photodiodes are thermally coupled to a heat sink which is heated to above ambient temperature in order that the temperature of the diodes may be closely controlled. During the course of a scan, the photosensitive diodes heat up as a result of the energy dissipated within the diodes during the scan interval. The resulting increase in temperature produces an increase in the output current for the photodiode for a fixed incident luminescent light intensity. To compensate for this change, the temperature of the diodes is measured before the beginning of a scan and at the end of the scan and the data acquired during that scan is compensated by subtracting a compensation signal from each data value. The compensation signal is determined on the basis of a linear increase in temperature during the scan period in combination with the scan time at which the data was taken.
With the desire for further increases in system linearity and decreases in system noise, improved photodetector temperature, linearity and noise control is needed.